Intraluminar perforated radially expandable drug delivery prosthesis and a method for the production thereof

ABSTRACT

A radially expandable prosthesis for implantation in a lumen comprises a tubular wall having an inner surface and an outer surface. The tubular wall is provided with cuts to form solid struts which have a thickness and which enables the prosthesis to expand. The solid struts have reservoirs made therethrough in the form of perforating holes for containing a therapeutic agent. The perforating holes each have an inner opening and an outer opening of substantially the same size. The prosthesis, including said perforating holes, has a smooth electrochemically polished surface.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a Continuation application from U.S. patentapplication Ser. No. 11/353,601, filed Feb. 14, 2006, which is aContinuation application from U.S. patent application Ser. No.09/798,990, now U.S. Pat. No. 7,135,039, filed Mar. 6, 2001, and whichclaims priority to European Patent Application 00870035, filed Mar. 6,2000, the entire contents of each being incorporated herein byreference.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH

Not Applicable

BACKGROUND OF THE INVENTION Field of the Invention

The present invention relates to radially expandable prostheses forimplantation in a lumen comprising a tubular wall produced from sheetmetal and showing cuts enabling the prosthesis to expand and to a methodfor producing such prostheses wherein said cuts are at least partiallymade by means of a laser beam.

In practice, intraluminal prostheses are generally known. They can beimplanted in a lumen, for example an artery, to strengthen, support orrepair the lumen. With coronary balloon dilatation for example, often aprosthesis is implanted in the place where a coronary artery is injuredor where it tends to collapse. Once implanted, the prosthesisstrengthens that part of the artery in a way the blood flow is ensured.A prosthesis configuration which is extremely suited for implantation ina body lumen, is a generally cylindrical prosthesis which can radiallyexpand from a first small diameter to a second larger one. Suchprostheses can be implanted in the artery by placing them on a catheterand transporting them through the artery to the desired location. Thecatheter is provided with a balloon or another expansion mechanism whichexerts a radial outwards pressure on the prosthesis so that theprosthesis expands to a larger diameter. These prostheses aresufficiently strong to stay in shape after expansion, even after removalof the catheter.

Radially expandable prostheses are available in a variety ofconfigurations, in this way an optimal efficacy is ensured in differentparticular situations. The patents of Lau (U.S. Pat. Nos. 5,514,154,5,421,955, and 5,242,399), Baracci (U.S. Pat. No. 5,531,741), Gaterud(U.S. Pat. No. 85,522,882), Gianturco (U.S. Pat. Nos. 5,507,771 and5,314,444), Termin (U.S. Pat. No. 5,496,277), Lane (U.S. Pat. No.5,494,029), Maeda (U.S. Pat. No. 5,507,767), Marin (U.S. Pat. No.5,443,477), Khosravi (U.S. Pat. No. 5,441,515), Jessen (U.S. Pat. No.5,425,739), Hickle (U.S. Pat. No. 5,139,480), Schatz (U.S. Pat. No.5,195,984), Fordenbacher (U.S. Pat. No. 5,549,662) and Wiktor (U.S. Pat.No. 5,133,732) all contain a sort of radially expandable prosthesis forimplantation in a body lumen.

The mentioned intraluminal prostheses have some disadvantages. Many ofthese expandable prostheses are not extremely flexible and have acentral axis that remains rather linear when the prosthesis is notexpanded yet. Due to such a lack of flexibility the insertion of theprosthesis in the artery to be correctly placed in the body lumen ishampered. Another problem of the intraluminal prostheses is theirdecrease in axial length at radial expansion. Although the patent of Lau(U.S. Pat. No. 5,514,154) attempts to reduce the axial shortening, itfails to succeed entirely.

When a prosthesis is placed in the artery or in another lumen, theimplantation has to be performed precisely in the desired place.Intraluminal prostheses are often exactly placed before their expansion,but due to the expansion the axial shortening causes that the prosthesisfinally does not turn up in the correct place.

In addition the determination of the exact location of a prosthesisduring an implantation in a lumen is difficult, although a highlyqualitative medical monitoring system is available. The problem of theexact place determination of the prosthesis enlarges the problem of theprecise and exact placement. There exists a need for a radiallyexpandable prosthesis presenting little or none axial shortening atradial expansion and that can be located without difficulty usingmedical imaging systems during the implantation. Another frequentlyoccurring problem is the occlusion of the side branches. In the case ofcoronary arteries this can cause a myocardial infarction.

Another important problem is the insufficient hemocompatibility ofintraluminal prostheses, when they are implanted intravascularly. Theycan cause acute or subacute thrombotic occlusions due to thrombusformation resulting in a considerable morbidity and even mortality.Furthermore these prostheses evoke a foreign body reaction with aconsiderable inflammation all around the prosthesis inducingfibromuscular cellular proliferation and narrowing of the prosthesis.The general object of the present invention is therefore to provide newmethods and new prostheses which enable to reduce the foreign bodyreaction against the implanted prostheses.

Concerning the method of producing such prostheses, EP-A-0 931 520teaches to start from a thin-walled tubular member, in particular astainless steel tubing, and to cut this tubing to remove portions of thetubing in the desired pattern for the prosthesis or stent. This cuttingprocess is performed by means of a computer-controlled laser. In orderto minimize the heat input by the laser into the prosthesis so as toprevent thermal distortion and other damages to the metal, use is madeof a Q-switched Nd/YAG laser which is operated to produce very shortpulses (<100 nsec) at a high pulse rate of up to 40 kHz. Further, a gasjet is created co-axially to the laser beam. Notwithstanding the use ofa gas jet, a considerable amount of debris, slag or molten material isformed along the edges of the cut which must be removed mechanically orchemically after the cutting operation. This is achieved in EP-A-0 931520 by soaking the cut stainless tube first for eight minutes in asolution of hydrochloric acid (HCl) and by subsequently electropolishingit in an acidic aqueous solution of sulfuric acid, carboxylic acids,phosphates, corrosion inhibitors and a biodegradable surface activeagent with a current density of about 0.06 to 0.23 amps per cm². Adrawback of such severe chemical and electrochemical polishing processesis that the inner and outer surfaces of the tubular prosthesis may alsobecome attacked.

The art referred to and/or described above is not intended to constitutean admission that any patent, publication or other information referredto herein is “prior art” with respect to this invention. In addition,this section should not be construed to mean that a search has been madeor that no other pertinent information as defined in 37 C.F.R. §1.56(a)exists.

All US patents and applications and all other published documentsmentioned anywhere in this application are incorporated herein byreference in their entirety.

Without limiting the scope of the invention a brief summary of some ofthe claimed embodiments of the invention is set forth below. Additionaldetails of the summarized embodiments of the invention and/or additionalembodiments of the invention may be found in the Detailed Description ofthe Invention below.

A brief abstract of the technical disclosure in the specification isprovided as well only for the purposes of complying with 37 C.F.R. 1.72.The abstract is not intended to be used for interpreting the scope ofthe claims.

BRIEF SUMMARY OF THE INVENTION

A first object of the present invention is therefore to provide a newmethod for cutting the prosthesis by means of a laser which enables toachieve cleaner or finer cut edges after the cutting operation and tolimit the required polishing operations to achieve an optimalbiocompatibility.

For this purpose, the method according to a first aspect of theinvention is characterized in that for making at least a number of saidcuts, said laser beam is guided in a jet of liquid towards the sheetmetal.

According to the invention, it has been found that by using such atechnique, which is known as water-guided laser technology, it ispossible to achieve immediately cleaner cut edges and to reduce thethermal and structural distortion, in particular the deformation of themetal grain structure. In this way, the body foreign reaction againstthe implanted prosthesis causing narrowing to the prosthesis can thus bemore easily avoided or reduced.

A second object of the present invention is to provide specific methodsfor electropolishing prostheses made from tantalum or from a nickeltitanium alloy which also enable to improve the biocompatibility of theprosthesis.

To this end, the method according to a second aspect of the invention ischaracterized in that when the prosthesis is made from a nickel titaniumalloy, in particular from nitinol, it is polished electrochemically inan electrolyte solution containing perchloric acid and at least onecarboxylic acid, in particular acetic acid and, when the prosthesis ismade from tantalum, it is polished electrochemically in an electrolytesolution containing sulfuric acid, hydrofluoric acid and optionally acarboxylic acid, in particular acetic acid.

In this second aspect of the invention, the prosthesis is also cut bymeans of a laser beam, preferably by means of the liquid guided laserbeam according to the first aspect of the invention. By using liquid, inparticular water-guided laser technology and the specificelectrochemical polishing techniques this prosthesis can be rendered ofa superior biocompatibility, which results in a reduced probability ofthrombogenicity, neointimal hyperplasia and intraluminal narrowing ofthe prosthesis after intraluminal implantation.

Further improvement of the biocompatibility of the prosthesis can beobtained according to the invention by applying a titanium nitride layerhaving a thickness of between 0.1 and 500 μm, and preferably a thicknessof between 1 and 10 μm.

In a third aspect, the invention relates to a special configurationwhich enables the prosthesis to release an effective amount oftherapeutic agent or medicine over a prolonged period of time, inparticular a medicine suppressing the foreign body reaction against theprosthesis increasing thereby also the biocompatibility of theprosthesis. The tubular wall of the prosthesis is provided with cutsforming struts having a predetermined thickness and enabling theprosthesis to expand, the struts having a longitudinal direction andshowing reservoirs made in said outer surface for containing thetherapeutic agent.

Such a prosthesis is already disclosed in EP-A-0 950 386. In this knownprosthesis, the reservoirs are formed by relatively shallow channelswhich are laser cut in the outer surface of the prosthesis. A drawbackof this known prosthesis is that at the location of the channels thelocal drug delivery will be much greater than at other locationsresulting in a quite non-homogeneous distribution of the therapeuticagent. Another drawback is that the depth of the channels is limited inview of the fact that the presence of the channels have a considerableeffect on the radial strength and durability of the prosthesis. Due tothe limited depth, the effect of this depth on the period of drugrelease is consequently also limited.

In a third aspect an object of the present invention is therefore toprovide a new prosthesis which enables to provide a more uniform drugrelease, to extend this release over a greater period of time and toincorporate the therapeutic agent in the prosthesis with a smallereffect on the radial strength thereof.

For this purpose, the prosthesis is characterized in this third aspectof the invention in that at least a number of said reservoirs are formedby holes which show at least an outer opening at the outer surface ofthe tubular wall and which extend over a depth in said struts largerthan 30%, preferably larger than 50% and most preferably larger than60%, of the thickness thereof, said outer opening having a widthmeasured perpendicular to said longitudinal direction and a lengthmeasured in said longitudinal direction which comprises at the most fivetimes, preferably at the most three times, said width.

Since the length of the holes comprises at the most five times the widththereof, more holes can be provided in the outer surface of theprosthesis, i.e. at shorter mutual distances, so that a more homogenousdrug delivery is possible. A further advantage of such shorter holes isthat they can be made deeper without affecting the required radialstrength of the prosthesis. In this way, it is possible to incorporatemore therapeutic agent in the prosthesis and to increase the releaseperiod thereof due to the fact that a larger amount of therapeutic agentcan be contained in one hole relative to the surface area of the outeropening thereof through which the therapeutic agent is released. Thesmall holes, which may show a bottom or extend entirely through thestrut wherein they are made, allow to load the prosthesis with a dose ofmedicine up to a thousand times higher compared to a none perforatedprosthesis. In this way a more biocompatible intraluminal prosthesis canbe obtained which can also be used as a vehiculum for releasing and ordepositing medicines locally.

In a preferred embodiment, at least a bottom portion of said hole issubstantially conical, the hole having either a bottom or extendingthrough the strut forming in said inner surface of the tubular wall aninner opening.

An important advantage of this embodiment is that the holes can be madeeasily by laser cutting, in particular in accordance with the liquidguided laser cutting technique, by simply directing the laser beam tothe desired spot and cutting the hole without any further movement ofthe laser beam. The depth of the hole can then simply be controlled byadjusting the total amount of energy of the laser beam, i.e. the pulsewidth, the duration and the intensity thereof. When making perforatingholes, the diameter of the inner opening of the holes on the inner sideof the strut can be controlled in the same way, i.e. also by adjustingthe amount of energy used to make the hole by means of the laser beam.In other words, the amount of therapeutic agent released towards theinside of the prosthesis can be easily controlled by selecting thedesired diameter of the inner openings. The total amount of cuttingenergy can be increased until the inner opening is substantially aslarge as the outer opening.

These and other embodiments which characterize the invention are pointedout with particularity in the claims annexed hereto and forming a parthereof. However, for further understanding of the invention, itsadvantages and objectives obtained by its use, reference should be madeto the drawings which form a further part hereof and the accompanyingdescriptive matter, in which there is illustrated and described aembodiments of the invention.

BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWING(S)

Other particularities and advantages of the invention will becomeapparent from the following description of some particular embodimentsof the method and the prosthesis according to the present invention. Thereference numerals used in this description relate to the annexeddrawings wherein:

FIG. 1 is a top plan view on a tubular prosthesis which has been cut inits longitudinal direction and pressed into a flat sheet;

FIG. 2 shows on a larger scale a portion of the sheet illustrated inFIG. 1 and showing additionally the holes provided in the outer surfaceof the prosthesis;

FIG. 3 is a diagram setting forth the different steps followed for theelectrochemical polishing of nitinol samples;

FIG. 4 is a microscopic picture (680×) showing the roughness of the sidesurface of a laser cut of the as-received nitinol alloy stent made bymeans of a conventional laser;

FIG. 5 is a microscopic picture (655×) of the nitinol alloy stentillustrated in FIG. 4 after having been subjected to the electrochemicalpolishing process;

FIGS. 6 and 7 are microscopic pictures (326×) of the over polishedsurface of a nitinol alloy stent;

FIG. 8 is a microscopic picture (300×) of the bottom surface of anelectrochemically polished nitinol alloy stent without pickling;

FIG. 9 shows, on a larger scale, a schematic cross-sectional view alonglines VI-VI in FIG. 2, illustrating a perforating hole with asubstantially cylindrical shape;

FIG. 10 is a view similar to FIG. 7 but showing a perforatingsubstantially conical hole having a substantially circular outer openingand a smaller inner opening;

FIG. 11 is a view similar to FIG. 8 but the non-perforating conical holedoes not extend through the strut but shows a bottom;

FIGS. 12 and 13 are views similar to the views of FIGS. 10 and 11 butshowing another shape of holes;

FIG. 14 is a scanning electron microscope picture (156×) showing aperforating hole of the same type as illustrated in FIG. 9;

FIG. 15 is a scanning electron microscope picture (156×) showing aperforating hole of the same type as illustrated in FIG. 10; and

FIG. 16 is a scanning electron microscope picture (625×) showing anon-perforating hole, provided with a bottom, of the same type asillustrated in FIG. 11.

DETAILED DESCRIPTION OF THE INVENTION

While this invention may be embodied in many different forms, there aredescribed in detail herein specific preferred embodiments of theinvention. This description is an exemplification of the principles ofthe invention and is not intended to limit the invention to theparticular embodiments illustrated.

For the purposes of this disclosure, like reference numerals in thefigures shall refer to like features unless otherwise indicated.

1) Description of the Basic Design of the Endovascular Prosthesis

In general the present invention relates to radially expandableprostheses for implantation in a lumen which comprise a tubular wallproduced from sheet metal and showing cuts enabling the prosthesis toexpand. In the method according to the invention, these cuts are atleast partially made by means of a laser beam. The prosthesis is thusmade starting from a tubular member wherein cuts are made, or whereinusually portions are cut away, according to the design of theprosthesis. Instead of starting from a tubular member, use could also bemade of a flat sheet which is enrolled and welded together to form thetubular prosthesis. FIGS. 1 and 2 illustrate a preferred embodiment of aradially expandable prosthesis that presents little or none axialshortening at radial expansion. The prosthesis consists of filaments orstruts 1 describing the outline of a cylindrical contour. Eachprosthesis filament connects to a separate surface at right angles to acentral axis of the cylindrical contour of the prosthesis and parallelwith other surfaces of the adjacent filaments. The prosthesis can existof a variable amount of filaments which all constitute the prosthesis.At least two filaments are necessary, including a first and a secondending filament to determine the extremities of the prosthesis contour.

These filaments all show a waving contour in the shape of consecutiveomegas. Consequently each filament is composed of a number of turns withlowest points and tops zigzag crossing over the length of each filament.The lowest point is the most distant from the adjacent filament and thetop is the most closely situated to the adjacent filament. FIG. 1 showsa typical configuration with 12 turns, a number that can vary from 3 to36 turns. The size of each filament, provided as the distance betweenlowest point and top, changes when the prosthesis expands radially,mostly the size diminishes. In FIG. 1 a typical configuration is shownwith a distance of 1.5 mm between the lowest point and top, thisdistance however can vary from 0.5 to 5 mm.

The end filaments are attached to adjacent intermediate filaments bymeans of connecting parts in the shape of an omega that act as axialelements joining two adjacent filaments. Such connecting parts are alsoable to fasten together intermediate filaments. Each connecting part isattached to the adjacent filaments with a first connection point to theone end of the connecting piece and a second one to the other end. Bothconnecting points are situated in the tops of the filaments. Thus theconnecting points are bridging the distance/opening between adjacentfilaments with the interstice as maximal width. Not necessarily allperforations are bridged with axial connecting parts. Separate outlinedintermediate elements can be joined together by means of junctions thatare connected with the intermediate elements on locations distant of thelowest points. Depending on the flexibility needs of the prosthesis avariable number of tops can be provided with connecting parts that linkadjacent filaments. In case a higher flexibility is necessary, more topswill stay empty with only one connecting piece between two adjacentfilaments. The prosthesis is constructed as such that during gradualexpansion of the prosthesis the filament waves will in a first phasebecome somewhat larger and than gradually become shorter. To compensatefor this shortening the omega shaped interconnections will graduallyenlarge resulting in a less axial shortening during gradual expansion.

2) Water Guided Laser Technology to Cut a Metallic IntraluminalProsthesis

Laser cutting of metallic tubes for example 316L stainless steel,nitinol, tantalum tubes or any other metallic tube causes a considerableheat release at the cutting surface, that radiates through the material.The disadvantage is that the metal structure (grain structure) isdeformed and that the cutting surface becomes irregular and oxidized.This leads to a considerable foreign body reaction against the implantedprosthesis causing narrowing of the prosthesis.

Utilizing water guided laser technology (Laser-Microjet™ or comparablesystems described for example in U.S. Pat. No. 5,902,499) wherein thelaser beam is guided in a jet of liquid, in particular of water, to cutthe omega-prosthesis we were able to diminish these disadvantages.

1) The water jet continuously removes the burnt metal particles,resulting in a better cutting surface. The water or other liquid isgenerally ejected out of a nozzle at a pressure of between 20 and 500bars, more particularly at a pressure of between 100 and 300 bars andpreferably at a pressure of about 150 bars. For creating the water jet,a nozzle showing an opening having an inner diameter of 20 to 100 μm, inparticular an inner diameter of 40 to 60 μm, and more particularly aninner diameter of about 60 μm can be used resulting in a water jet of asimilar diameter wherein the laser beam is contained due to thedifference in refractive indices of the liquid and air.

2) Thanks to the continuous liquid or water-cooling, the heatpenetration is lower, causing quasi no deformation of the metal grainstructure.

As an example we provide here the specific conditions to apply the waterguided laser cutting system. Stainless steel tubes with a diameter of0.0625 inch (1.585 mm) with a wall thickness of 0.004 inch (0.1 mm) wereplaced on a continuous rotating holder and cut with a Haas laser at afrequency of 100 Hz, pulse duration of 0.15 ms, voltage 510 Volt, heador water nozzle diameter of 60 μm and water pressure of 150 bar.Comparing the samples with conventional laser cut stents, the surfacelooked much brighter and less blackened. SEM examination showed a muchmore regular surface. Implantation of these stents after degreasing,ultrasonic cleaning and sterilization in porcine coronary arteriesresulted in considerably less thrombus formation adjacent to the stentfilaments and a moderate inflammatory response at 6 days follow-up. At 6weeks follow-up area stenosis was 60%. These results however comparefavorably with conventional laser cut stents. Conventional laser cutstents implanted under the same conditions resulted in a totalthrombotic occlusion of the stented vessel in 40% of the cases, anabundant inflammatory response at 6 days follow-up and an area stenosisof 86% at 6 weeks follow-up. This water guided laser technology can beused to cut any other coronary stent or endovascular prosthesis out of ametallic tube.

3) Electrochemical Polishing of a Metallic Intraluminal Prosthesis

Surface characteristics of metal intraluminal prosthesis are determiningthe human foreign body response to the intraluminal prosthesis.Therefore optimal surface characteristics are critical for the acute andlate patency of an intraluminal prosthesis. To further optimize thecutting surface, specific electrochemical polishing techniques were usedto optimize the surface characteristics of intraluminal prostheses.Depending on the material used, specific chemical solutions weredeveloped for optimal electrochemical polishing of prostheses.

Basic Principles of Electropolishing

Electropolishing is a process by which metal is removed from a workpiece by passage of electric current when the work piece is immersed ina liquid media (electrolyte). The work piece is connected to the anodicterminal, while the cathodic terminal is connected to a suitableconductor. Both anodic and cathodic terminals are submerged in thesolution, forming a complete electrical circuit. The current applied isdirect (DC) current. In this process, the work piece is dissolved,adding metal ions to the solution. When a current passes through theelectrolyte, a liquid layer of anodic dissolution products is formed onthe surface of the anode; this layer has a higher viscosity and greaterelectrical resistivity than the bulk of the electrolyte. The thicknessof the liquid layer on a rough surface differs from site to site. Thecurrent density is non-uniform as result of such non-uniform liquidlayer; i.e. it is higher on peaks than in crevices. Thus, peaks dissolvemore rapidly than crevices, this, therefore, produces a surface-levelingeffect.

Furthermore electrochemical polishing results in a superficial oxidelayer (passivation) which plays also an important role in thebiocompatibilisation of a foreign body.

The quantity of metal removed from the work piece is mainly proportionalto the amount of current applied and the time during which the currentis applied. In addition, the geometry of the work piece can affect thedistribution of the current and, consequently, has an important bearingupon the amount of the metal removed in local area.

Factors Affecting the Electropolishing Process

The mode of anodic dissolution of a metal may depend on its nature, thesurface state, the composition of the electrolyte, and the temperature,current density and stirring during the electropolishing process.

Electrolytes used in electropolishing should satisfy the followingrequirements:

1) high-quality polishing at low voltages and current densities,

2) wide working range of anodic current densities and temperature,

3) a high stability (during operation and upon storage) and long servicelife,

4) absence of attack on the metal when current does not flow,

5) electrolyte should consist of cheap, readily available materials andshould not present any safety hazards,

6) recovery after a certain period of service should be simple, e.g. byadditions of necessary components,

7) the throwing power of the bath should be good, i.e. samples ofcomplex shape should dissolve uniformly over their entire surface,

8) the ohmic resistivity should be low, i.e. the required currentdensity should be obtained at a low voltage,

9) the electrolyte should be suitable for use in the electropolishing ofmany metals.

The anodic potential, the anodic current density, and the appliedvoltage are the main electrical parameters of the electropolishingprocess. The process is controlled on the basis of the anodic currentdensity and sometimes on the basis of the applied voltage. For any metaland electrolyte system, there should be a certain optimal anodic currentdensity, which provides the highest-quality electropolishing.

The temperature of the electrolyte has a marked effect on the polishingquality. A range of optimum temperatures should exist for anymetal-electrolyte system. A drop in the temperature increases theviscosity of the electrolyte and thus reduces the rate of diffusion ofanodic dissolution products from the anode surface to the bulk of theelectrolyte.

The electropolishing time should decrease with increasing currentdensity or with decreasing initial roughness of the surface. The initialroughness and the state of the surface also affect electropolishingquality. Before electropolishing, the surfaces are preferably degreasedand cleaned in organic solvents or by chemical etching in suitablesolutions. Stirring is used in cases that the anode is coated with somesoluble films or it is necessary to remove bubbles adhering to thesurface. Stirring of the electrolyte requires an increase in the currentdensity. The cathodes used in electropolishing should not be attacked inthe polishing solution. The surface area of the cathode is preferablymuch greater than the surface area of the polished work piece. Thisensues a more uniform current distribution, reduces cathodicpolarization and reduces power losses. After the electropolishing, thework pieces should be washed with water or other solvents in order toremove residues of the electrolyte or the anodic dissolution products.

Description of Electrochemical Polishing Techniques for Nitinol andTantalum Endoluminal Prosthesis

The electropolishing device that we used was self-designed. A glasscontainer (150 ml) was used as container for the electrolyte. A DCrectifier (Polipower, Struers, Denmark) was employed as a power supply.A nitinol sheet material (length 15 cm, width 2.5 cm and thickness 0.2cm) was selected as anode. As shown in FIG. 3, the as-received sampleswere first cleaned with an alkaline solvent with a detergent additive inan ultrasonic bath for more than ten minutes. They were then cleaned indistilled water with an ultrasonic agitation device for more than tenminutes. Since the sheet materials were covered by a black oxide film,they were pickled at room temperature in an acid solution as follows: 2ml hydrofluoric acid (38-40%) and 40 ml nitric acid (14 M) for differenttime durations in a glass container (50 ml). By observation of thechange of the surface state of these sheet samples, a time of sevenminutes was finally chosen because the black film just disappeared atthis time duration. After pickling, the samples were rinsed in distilledwater with an ultrasonic agitation device for more than 10 minutes.After the preparation processes above, electropolishing was then studiedwith several selected electrolyte mixtures shown in Table 1. Solutions(i) and (ii) were used for electropolishing. Chemical polishing was alsoevaluated using the solutions (iii) and (iv) in order to have acomparison of the chemical polishing effect with electrochemicalpolishing.

TABLE 1 Selected mixtures for polishing nitinol alloy sheet materialsSolution (% concentration) Volumes i) perchloric acid (70%) 6 ml aceticacid (99.8%) 94 ml (ii) perchloric acid (70%) 5 ml acetic acid (99.8%)100 ml (iii) H₂O₂ 50 ml HF (48-51%) 5 ml (iv) H₂O₂ 75 ml HF (48-51%) 5ml

Several conditions were selected for the different polishing mixtures inorder to compare the differences in the effect of polishing and then tooptimize the condition and effect of polishing, which is shown in Table2. Firstly, as shown in Table 2 (a), the electrolyte (i) was used with afixed applied voltage for different times to test the electropolishingconditions. The effects were evaluated visually and with opticalmicroscopy. Secondly, the conditions of fixed time and applied voltagewere studied as well as that of some other selected times (Table 2 (b)).Then, the electrolyte (ii) was used and the conditions were changedsimilar to those of electrolyte (i) (Table 2 (c) and Table 2 (d)). Thesamples were also immersed in different mixtures of acids for differenttimes, as shown in Table 2 (e).

TABLE 2 The processing conditions for polishing nitinol alloy sheetmaterial Table 2 (a): Applied Voltage Anodic current Time Electrolyte(V) (amp) (min.) Electrolyte (i) 30 0.2~0.3 2 3 4~5 6 8 Table 2 (b):Applied Voltage Anodic current Time Electrolyte (V) (amp) (min.)Electrolyte (i) 5 0.023 10 10 0.045 10 15 0.1 10 and 15 20 0.16 3, 15and 10 25 0.22 3, 5 and 10 Table 2 (c): Applied Voltage Anodic currentTime Electrolyte (V) (amp) (min.) Electrolyte (ii) 30 0.15~0.25 2 3 4~56 8 10  Table 2 (d): Applied Voltage Anodic current Time Electrolyte (V)(amp) (min.) Electrolyte (ii) 20 0.15 3 6 25 0.17 3 6 Table 2 (e): TimeElectrolyte (min.) Electrolyte (iii) 3, 9, and 15 Electrolyte (iv) 3, 9,and 15

Most of the polishing processes were conducted at room temperaturewithout stirring. A part of the polishing processes however was done atan elevated temperature. Also the effect of stirring on the polishingprocess was explored. The reason that the applied voltage was selectedas one of the controlling parameters was that the current was notstationary during the process of electropolishing and the stents had aspecial shape of mesh so that the current density was very difficult tocalculate accurately. All of the polished samples were rinsed indistilled water with an ultrasonic agitation device for more than 10minutes and then stored in ethanol. Evaluation of the effects of thepolishing was performed by means of optical microscopy and scanningelectron microscopy.

Electropolishing of a Nitinol Endovascular Prosthesis, for Example aCoronary Stent

The same electropolishing cell as that for the sheet material wasemployed. First of all, due to the finite stent samples and their highcost, electropolishing of some nitinol alloy wires of differentdiameters of 1 mm, 0.3 mm and 0.5 mm were performed with differentparameters in order to find an optimal way of electropolishing thenitinol alloy stents. All the wires were covered with a black oxidelayer. They were ground with rough abrasive paper to remove the oxidelayer. The different conditions for the electropolishing are shown inTable 3.

TABLE 3 The selected processing conditions for electropolishing nitinolalloy wires Applied Voltage Anodic current Time Electrolyte (V) (amp)(min.) Electrolyte (ii) 30 0.21 2, 3, 4~5, 8 25 0.2 2, 3, 4~5, 8 20 0.182, 3, 4~5, 8

The selection of these conditions was based on our observations of theelectropolishing of sheet materials. The removal was measured with amicrometer (Mitutoyo Digimatic micrometer).

The as-received stents, made by means of a conventional laser cuttingprocess and illustrated in FIG. 4, were first cleaned with an alkalinesolvent with detergent additive in an ultrasonic bath for more than tenminutes in order to remove the contaminants of the surface. All thesamples were then cleaned in distilled water with an ultrasonicagitation device for more than ten minutes. Considering the oxide layerformed during the process of fabrication, the stents were pickled for 2,4 and 6 minutes at room temperature in the following acid solution: 2 mlhydrofluoric acid (38-40%) and 40 ml nitric acid (14 M).

By observation with optical microscopy, a time of nearly six minutes waschosen as the optimum pickling time of the stents. The samples were thenrinsed in distilled water with ultrasonic agitation for more than 10minutes. After pickling and rinsing, electropolishing was done withoutstirring at room temperature using the conditions shown in Table 4.

TABLE 4 The selected processing conditions for electropolishing nitinolalloy stents Applied Voltage Anodic current Time Electrolyte (V) (amp)(min.) Electrolyte (ii) 30 0.25 3 25 0.17 1, 1.5, 3 20 0.15 1, 1.5~2

These conditions were selected according to the results of theelectropolishing of sheet materials and wires, considering the specificthin shape of mesh of the stents. After electropolishing, the sampleswere rinsed in distilled water with an ultrasonic agitation device formore than 10 minutes. A scanning electron microscopic picture of aelectropolished stent is given in FIG. 5 illustrating the much smoothersurface compared to the stent prior to electropolishing.Electropolishing of the stents without the pre-treatment of pickling wasalso done in order to check whether or not the oxide layer can beremoved as well as to investigate the effects of pickling on theelectrochemical polishing of the stents for the following condition(Table 5).

TABLE 5 The conditions for electropolishing the stents without picklingApplied Voltage Anodic current Time Electrolyte (V) (amp) (min.)Electrolyte (ii) 20 0.15 1.5~2

Table 6 summarizes all the results of applied polishing processesprocess for nitinol alloy materials.

TABLE 6 The comparison of the different polishing process Table 6 (a):Applied Voltage Anodic current Time Process Material Electrolyte (V)(amp) (min.) Result Electropolishing Sheet Electrolyte (i)* 30 0.2~0.3 2less-polished 3 general 4~5 good 6 overpolished 8 overpolished 5 ~0.02310 no changes 10 ~0.045 no changes 15 ~0.1 small changes 20 ~0.16general 25 ~0.22 general 15 ~0.1 15 attacked 20 ~0.16 3 general 25 ~0.22general 20 ~0.16 5 good 25 ~0.22 good Table 6 (b): Applied VoltageAnodic current Time Process Material Electrolyte (V) (amp) (min.) ResultElectropolishing Sheet Electrolyte (ii)** 30 ~0.2 2 general 3 good 4~5better 6 general 8 overpolished 20 ~0.15 3 general 25 ~0.17 general 20~0.15 6 good 25 ~0.17 good Chemical Sheet Electrolyte (iii)# — — 3 nochanges polishing — — 9 rough — — 15 rough Electrolyte (iv)## — — 3 nochanges — — 9 rough — — 15 rough Table 6 (c): Applied Voltage Anodiccurrent Time Process Material Electrolyte (V) (amp) (min.) ResultElectropolishing Wire Electrolyte (ii) 30 ~0.21 2 general 3 good 4~5overpolished 8 overpolished 25 ~0.2 2 general 3 good 4~5 overpolished 8overpolished 20 ~0.18 2 general 3 good 4~5 overpolished 8 overpolishedTable 6 (d): Applied Voltage Anodic current Time Process MaterialElectrolyte (V) (amp) (min.) Result Electropolishing Stent Electrolyte(ii) 30 ~0.25 3 overpolished 25 ~0.17 1 general 1.5 general 3overpolished 20 ~0.15 1 good 1.5~2 better *perchloric acid (70%) 6 ml,acetic acid (99.8%) 94 ml **perchloric acid (70%) 5 ml; acetic acid(99.8%) 100 ml #H₂0₂ 50 ml, HF 5 ml ##H₂0₂ 75 ml, HF 5 ml

All these studies were conducted at room temperature without agitation.Several elevated temperatures (25° C. and 30° C.) were used, but nolarge changes on the results of polishing were found. As to theagitation, it gave relatively bad results for polishing sheet materials.During observation of the polishing process, dark spots appeared on thesurface, and by means of optical microscopy, these surfaces were provento be rough. In addition, bubbles were found adhering to the surface ofthe sample when polishing.

Two electrolytes were selected:

(i) perchloric acid (70%) 6 ml, acetic acid (99.8%) 94 ml

(ii) perchloric acid (70%) 5 ml, acetic acid (99.8%) 100 ml

Finally electrolyte (ii) was selected for electrochemical polishing ofeither sheet materials or stents (Table 1). During experiments withelectrolyte (i), no change on the surface of the sheet material wasfound using voltages of 5 V and 10 V, even for more than ten minutes.This is consistent with the fact that electrochemical polishing takesplace only when the current density is higher than that noted at thecritical point. Current density and voltage are closely related inpolishing. As voltage increases there is an increase in current densitygenerally. The surface is attacked with a voltage of 15 V for 15minutes. This might be the effect of electroetching. Electroetching is acomparatively slow process and its current densities are often smallerthan those with electropolishing. Thus, it is suggested to select thevoltages of 20 V, 25 V and 30 V for this nitinol alloy material from theresults in Table 6.

The duration of polishing for sheet materials is longer than that ofwires and much longer compared to that of stents. This might be becauseof the difference in degree of surface roughness of these materials andrelated to the thin size of the stent filaments. A time, up to six oreight minutes is so long that it causes over polishing for the sheetmaterials. However, polishing time of three minutes has caused overpolishing for the stents. FIGS. 6 and 7 show two overpolished surfacesof stents. One is apparently attacked, which shows a very bad qualityand the size of the stent did not remain uniform. Another one has arelatively smooth and uniform surface but the amount removed might be solarge that the stent has been apparently too thin to have enoughstrength to be used. Thus, the amount removed from stents should becontrolled very carefully so that the mechanical strength of the stentis maintained while the smoothness is obtained by means ofelectrochemical polishing.

The smooth surface can not be obtained by means of chemical polishingalone in this experiment. Chemical polishing is not sufficient to polishthese nitinol alloy materials. The optimal conditions forelectrochemical polishing of nitinol stents are shown in Table 7. Theseoptimal conditions will be somewhat different for each particularnitinol sample and will have to be restudied for each particular nitinolprosthesis depending on the design and mesh thickness used.

TABLE 7 The optimal condition for electropolishing of nitinol alloystents Applied Voltage Anodic current Time Electrolyte (V) (amp) (min.)Electrolyte (ii)** 20 0.15 1.5~2 **perchloric acid (70%) 5 ml, aceticacid (99.8%) 100 ml

Based on the performed tests, the electrolyte solution used forelectrochemically polishing nickel titanium alloys, in particularnitinol, comprises at least perchloric acid and at least one carboxylicacid, in particular acetic acid. In a preferred embodiment, theelectrolyte solution comprises the perchloric acid and the acetic acidin a concentration corresponding to the perchloric acid and acetic acidconcentrations in a mixture of 1 to 10 volume parts of a 70% by weightperchloric acid solution and 2 to 500 volume parts of a 99.8% by weightacetic acid solution, and more particularly in a concentrationcorresponding to the perchloric acid and acetic acid concentrations in amixture of about 5 to about 6 volume parts, preferably about 5 volumeparts, of a 70% by weight perchloric acid solution and about 94 to about100 volume parts, preferably about 100 volume parts, of a 99.8% byweight acetic acid solution. As explained hereabove, the most preferredelectrolyte solution comprises 5 volume parts of a 70% perchloric acidsolution and 100 volume parts of a 99.8% acetic acid solution.

Influence of the Preparation with Acidic Pickling on ElectrochemicalPolishing

The conditions and results of electrochemical polishing stents with nopreparation are summarized in Table 8.

TABLE 8 The result of electropolishing stents with no pickling AppliedVoltage Anodic current Time Process Material Electrolyte (V) (amp)(min.) Result Electropolishing stent Electrolyte (ii) 20 ~0.15 1.5~2 Bad

FIG. 8 shows the morphology of the stent surface electropolished withoutacidic pickling. It is very clear that the rough oxide layers stilladhere to the bottom surface of the stent and reveal an even worsequality than the as-received one. Thus, it can be concluded that in casethe stent is covered by a heavy oxide layer such a layer can not beremoved by means of electrochemical polishing alone, i.e. thepreparation of pickling is necessary for electrochemical polishing ofnitinol alloy stents which are covered with heavy oxide layers.

In this experiment, acidic pickling was explored for stents at roomtemperature with acidic solution: 2 ml hydrofluoric acid (38-40%) and 40ml nitric acid (14 M).

Several periods of time were selected. The description of the effects ofthis pickling process is summarized in Table 9.

TABLE 9 The effects of acidic pickling for preparation ofelectropolishing stents Time (min.) 1 2 4 6 8 Result no change no changegeneral good overpickled

During the immersion times 1 min and 2 min., there was no apparentchange on the surfaces compared with that of the as-received sampleswhen studied by optical microscopy. For 4 min., 6 min. and 8 min. theoxide layers were removed. There were still some oxides adhered to thesurface immersed for 4 min., whereas the oxide layer was removedcompletely for incubation times of 6 min. and 8 min. Thus, a time of 5to 7 min. was selected as optimal pickling time for nitinol stents.

Polishing of a Tantalum Intraluminal Prosthesis, for Example a CoronaryStent

This study was done by means of electrochemical polishing and chemicalpolishing in order to find an optimal condition of polishing of tantalumstents and to obtain a better quality of the surface of tantalum stents.The as-received materials were non-polished tantalum stents. Thepolishing cell was designed similar to that for nitinol alloy materialmentioned before. A glass container (100 ml) was used as an electrolytecontainer. A DC rectifier (Polipower Struers) was employed as a powersupply. For cathode, a graphite stick was chosen with a diameter of 10mm. Several electrolytes were selected for this study, as shown in Table10.

TABLE 10 The selected electrolytes for polishing tantalum stentsSolution (% concentration) Volumes (I) acetic acid (99.8%) 20 ml H₂SO₄(95-97%) 50 ml HF (48-51%) 10 ml (II) H₂SO₄ (95-97%) 90 ml HF (48-51%)10 ml (III) H₂SO₄ (95-97%) 50 ml HNO₃ (65%) 20 ml HF (48-51%) 20 ml

The electrolytes (I) and (II) were used for electrochemical polishing,and the electrolyte (III) was used for chemical polishing. Theas-received samples were first cleaned with an alkaline solvent withdetergent additive dipped in an ultrasonic bath for more than tenminutes. All the samples were then cleaned in distilled water with anultrasonic agitation device for more than ten minutes. After degreasing,the samples were treated by means of acidic pickling for several periodsof time: 2.5, 5, 7.5, 10 and 20 minutes in the following solution: HF48-51% 5.6 ml, H₂SO₄ 95-97% 1 ml, HNO₃ 65% 8 ml, H₂O 8 ml. The sampleswere then cleaned in distilled water with an ultrasonic agitation devicefor more than ten minutes. According to the effects observed by means ofoptical microscopy, ten minutes was chosen as the best pickling time.After pickling, the samples were polished with the selectedelectrolytes. The conditions are given in Table 11.

TABLE 11 The conditions for polishing tantalum stents Applied voltageTime Electrolyte (V) (min.) Electrolyte (I) 2.5, 5, 7.5, 10 2  5 0.5, 1,2, 3 Electrolyte (II) 15 3, 6, 9, 12 10, 15, 20 9 Electrolyte (III) 2,4, 6

The voltage was selected as the controlling parameter because thecurrent density was difficult to determine with the specific shape ofthe stents. The electropolishing of the as-received samples withoutdegreasing and pickling was also done in order to check whether or notan oxide layer exists on the surface as well as to investigate theeffects of pickling on the electrochemical polishing of the stents. Theconditions are shown in Table 12.

TABLE 12 The conditions and the result of electropolishing tantalumstents without degreasing and pickling Applied voltage Time Electrolyte(V) (min.) Result Electrolyte (I) 5 2 Bad Electrolyte (II) 15 9 Bad

After cleaning in distilled water with an ultrasonic agitation devicefor more than ten minutes, all the polished samples were evaluated withoptical microscopy and some of them were then studied by means ofscanning electron microscopy. The SEM pictures of both the polishedsamples and the as-received samples were taken in order to compare theirsurface qualities.

Comparison Among the Different Polishing Methods

Table 13 summarizes all the results of the explored polishing processesfor the tantalum stents. All these electrochemical polishing processeswere conducted at room temperature with agitation. Prior to theelectropolishing processes, both degreasing and pickling were done. Twoelectrolytes were explored for electropolishing: (I) acetic acid 20 ml,H₂SO₄ 50 ml, HF 10 ml, (II) H₂SO₄ 90 ml, HF 10 ml.

TABLE 13 The comparison of the different polishing processes AppliedVoltage Time Process Electrolyte (V) (min.) Result ElectropolishingElectrolyte 2.5 2 general (I) 5 better 7.5 good 10 overpolished 5 0.5less polished 1 general 2 good 3 overpolished Electrolyte 3 lesspolished (II) 15 6 general 9 good 12 overpolished 10 less polished 15 9good 20 general Chemical Electrolyte 2 rough polishing (III) 4 rough 6rough

From the evaluation of the effects of polishing by means of opticalmicroscopy, electrolyte (I) gave a relatively better effect; thereforeit was concluded to be the preferred electrolyte.

The voltages 2.5 V, 5 V, 7.5 V and 10 V were explored respectively withelectrolyte (I) for 2 min. The voltages of 2.5 and 10 V providedrelatively bad results. 5 V revealed the best results among thesevoltages. Fixing the voltage to 5 V, the times 0.5 min., 1 min. and 3min. were selected in order to compare the results. The result was thatthe voltage 5 V and the time 2 min. in conjunction with electrolyte (I)were optimal parameters for electrochemical polishing this kind oftantalum stents. Electrolyte (II) was also used with some changedparameters. Several times were explored with fixed voltage 15 V. It wasfound that 9 min. causes the best result among these times, and either 3min. or 12 min. resulted in a bad surface quality. Then the time 9 min.was fixed and voltages were changed to explore their effects on thepolishing. 15 V gave a relatively better result than 10 V and 20 V.Similar to the results in the experiment of polishing nitinol alloymaterials, chemical polishing could not lead to a sufficient smoothsurface. The conditions and the results were summarized in Table 13.

Based on the performed tests, the electrolyte solution used forelectrochemically polishing tantalum comprises at least sulfuric acidand hydrofluoric acid, and optionally at least one carboxylic acid, inparticular acetic acid. In a preferred embodiment, the electrolytesolution comprises sulfuric acid and hydrofluoric acid in aconcentration corresponding to the sulfuric acid and hydrofluoric acidconcentrations in a mixture of 70 to 120 volume parts of a 95 to 97% byweight sulfuric acid solution and 5 to 20 volume parts of a 48 to 51% byweight hydrofluoric acid solution, and more particularly in aconcentration corresponding to the sulfuric acid and hydrofluoric acidconcentrations in a mixture of about 90 volume parts of a 95 to 97% byweight sulfuric acid solution and about 10 volume parts of a 48 to 51%by weight hydrofluoric acid solution. In a further preferred embodiment,the electrolyte solution comprises sulfuric acid, hydrofluoric acid andacetic acid in a concentration corresponding to the sulfuric acid,hydrofluoric acid and acetic acid concentrations in a mixture of 20 to80 volume parts of a 95 to 97% by weight sulfuric acid solution, 5 to 20volume parts of a 48 to 51% by weight hydrofluoric acid solution and 10to 30 volume parts of a 99.8% by weight acetic acid solution, and moreparticularly in a concentration corresponding to the sulfuric acid,hydrofluoric acid and acetic acid concentrations in a mixture of about50 volume parts of a 95 to 97% by weight sulfuric acid solution, about10 volume parts of a 48 to 51% by weight hydrofluoric acid solution andabout 20 volume parts of a 99.8% by weight acetic acid solution.

Influence of Sample Preparation with Degreasing and Pickling onElectrochemical Polishing

The conditions and the results of electrochemical polishing with nopickling were summarized in Table 12. SEM evaluation of the stentsurface electrochemically polished without degreasing and acidicpickling showed disappointing results. It was clear that the rough oxidelayers still adhered to the side surface of the stent. The surface ofthe polished sample reveals a worse quality than that of the as-receivedone. Thus, it can be concluded that the heavy oxide layer can not beremoved only by means of electrochemical polishing, i.e. the preparationof pickling is necessary before electrochemical polishing the tantalumstents which are covered with heavy rough oxide layers. Degreasing wasaccomplished by an alkaline solvent with detergent additive. Picklingwas used to remove the heavy oxide layer and normally done with alkalineor acidic solutions. In this experiment, acidic pickling was explored atroom temperature with acidic solutions: HF 48-51% 5.6 ml, H₂SO₄ 95-97% 1ml, HNO₃ 65% 8 ml, H₂O 8 ml.

Several time periods were selected. The description of the effects ofthis pickling process was summarized in Table 14.

TABLE 14 The effects of acidic pickling for preparation ofelectropolishing stents Time (min.) 2.5 5 7.5 10 20 Result no change nochange general good overpickled

For the immersion times 2.5 min. and 5 min., there were no apparentchanges on the surface compared to that of the as-received samples byobservation with optical microscopy, whereas for a time more than 7.5min. some precipitates began to appear in the solutions during theexperiment. The immersion time 20 min. caused attack on the surface.Thus, the pickling time should be controlled seriously in order toremove all the oxide layers and to avoid surface attack. In this study 6min. was found to have a relatively good effect.

Electrochemical Polishing: Conclusions Nitinol

For nitinol a pre-treatment using a solution of 2 ml of hydrofluoricacid and 40 ml of nitric acid (14 M) for 5 to 7 minutes is suggested.For electrochemical polishing optimal results were found with 5 ml ofperchloric acid (70%) and 100 ml of acetic acid (99.8%) using an anodiccurrent of 0.15 amp and a voltage of 20 V during 1 to 3 minutes (Table7).

Tantalum

For tantalum a solution of 20 ml of acetic acid, 50 ml of hydrosulphateand 10 ml of hydrofluoride or a solution of 90 ml of hydrosulphate and10 ml of hydrofluoride was used. The voltage was 5 V during a period of1 to 5 minutes and the pre-treatment was done using a solution of 5.6 mlof hydrofluoride (48-51%), 1 ml of hydrosulphate (95-97%), 8 ml ofhydronitrate and 8 ml of water during 5 to 7 minutes (Table 15).

TABLE 15 The optimal conditions for electropolishing the tantalum stentsApplied voltage Time Electrolyte (V) (min.) Electrolyte (I)*** 5 ~2***acetic acid 20 ml, H₂SO₄ 50 ml, HF 10 ml

This treatment resulted in a further reduction of thrombogenicity, offoreign body reaction and of prosthesis narrowing. These inventions canbe used for any stent or endovascular prosthesis made of nitinol ortantalum.

4) Turning a Metallic Endovascular Prosthesis More Biocompatible Using aTitaniumnitride Coating

A metal prosthesis always causes a kind of foreign body reaction afterintraluminal implantation.

To improve the biocompatibility of the prosthesis, fine coatings can beapplied. In a preferred embodiment of the invention, the prostheses arecovered with a titanium nitride coating showing a thickness of between0.1 and 500 μm and preferably a thickness of between 1 and 10 μm.Experiments with titanium nitride coatings (1-15 μm) showed asignificantly decreased foreign body reaction in a porcine coronarymodel, what did result in a significant amelioration of the minimalluminal diameter of the prosthesis at follow-up.

Description of the Titanium Nitride Coating. Evaluation of the TinCoated Endovascular Prosthesis in a Porcine Coronary Model

Titanium nitride (TiN) coatings have proved their efficiency inincreasing the lifetime of cutting tools. Their tribological propertiesare widely known and their use in bioengineering applications as abiomaterial has been considered, particularly as a wear-resistantcoating for Ti6A14V orthopedic implants. The tests undertaken showedthat wear was reduced, that the TiN friction coefficient was low andthat TiN presented good chemical stability.

The present inventors used TiN (5 μm) to coat an intraluminal prosthesisand demonstrated improved biocompatibility.

To illustrate the invention a coronary stent of a coil-type design, asdescribed in U.S. Pat. No. 5,183,085 was used. It consisted of apreconditioned, non ferromagnetic, highly polished stainless steel wire(AISI 316L) with a diameter of 0.18 mm. This design allows folding(radial compression) on any conventional balloon, resulting in a lowprofile 6F guiding catheter compatible stent delivery system. Percentageof axial shortening upon expanding the balloon is less than 5% and thestent is available in lengths from 12 mm up to 40 mm allowing customizedstenting. These stents are available as bare stents or as mountedstents. In the present example stents of a length of 16 min were used.For this invention any laser cut stainless steel mesh stents or anyintraluminal metal prosthesis can be used as well.

Porous Tin Coating

The vacuum deposition techniques of physical vapor deposition (PVD) andchemical vapor deposition (CVD) are well known for their ability to formTiN coatings of different structures and stoichiometries. Porous TiN canbe generated in a reactive sputtering process, which is a special PVDmethod.

In the reactive sputtering process, Ar ions are produced by a glowdischarge and accelerated against a Ti target. The ions impinging on theTi lead to the ejection of particles from the target surface. Theseparticles condense on the surfaces in line-of-sight to the target.Additional N₂ gas activated in the plasma and the Ti react to yield TiN.The structure of the TiN can be influenced and determined by controllingthe deposition parameters like Ar and N₂ pressure, target power,substrate bias voltage, and substrate position relative to the Titarget. For certain parameters, the TiN layer grows with columnarstructure and shows up to 1000 fold increase in effective surface area.The thickness of the sputtered layer lies around 5 μm. The thicknessconstancy is ensured by a uniform rotation of the intraluminalprosthesis during the sputtering process.

Experimental Work

The TiN coated and bare non-coated stents were radially compressed on aballoon catheter (3 to 3.5 mm) and randomly implanted in a series ofcoronary arteries of 20 domestic cross bred pigs of both sexes, weighing25 to 30 kg. Ten TiN coated stents and 10 non-coated highly polishedstainless steel stents were implanted for comparison. All stentdeployments and implantations were successful and resulted in properlystented vessel segments. Six weeks after implantation, controlangiography of the stented vessels was performed and subsequently pigswere sacrificed. At that time their average weight was about 70 kg andthe vessels had also grown considerably, compared to their size at thetime of implantation.

Angiographic analysis (quantitative coronary angiography) of stentedvessel segments was performed before stenting, immediately afterstenting, and at follow-up using the Polytron 1000-system as describedby De Scheerder et al. in the Journal of Invasive Cardiology 1996;8:215-222. The lumen diameters of the vessels were measured before stentimplantation (=pre-stenting artery diameter values), immediatelythereafter (=post-stenting values) and at follow-up (=diameters after 6weeks). The degree of oversizing (%) was expressed as measured maximumballoon size minus minimal stent lumen diameter (measured 15 minutesafter stent implantation) and divided by measured maximum balloon size.The late loss value is an indication of hyperplasia and is thedifference between the post-stenting value and the diameter atfollow-up. The results of the angiographic measurements are summarizedin Table 16. Baseline selected arteries, measured balloon diameter andpost stenting diameter were similar for the three types. Oversizing andrecoil were also similar. At six weeks follow-up a larger minimalluminal stent diameter and a decreased late loss was found for theTiN-coated stents.

TABLE 16 Quantitative Coronary Analysis of titanium nitride coatedstents Mean Artery Non-coated stent TiN-coated stent Diameter (mm) n =10 n = 10 Pre stenting (mm) 2.52 ± 0.18 2.53 ± 0.27 Balloon size (mm)2.93 ± 0.16 2.94 ± 0.15 Post stenting (mm) 2.68 ± 0.16 2.71 ± 0.18Oversizing (%) 16 ± 6  16 ± 7  Recoil (%) 8 ± 4 8 ± 4 6 weeks FU (mm)2.52 ± 0.29 2.69 ± 0.24 Late loss 0.16 ± 0.28 0.02 ± 0.16

After the pigs were sacrificed, coronary segments were carefullydissected together with 10 mm minimum vessel segment both proximal anddistal to the stent. Histopathology, as evaluated by light microscopicexamination, was performed on very thin cross-section slices of thestented artery sections. Injury of the arterial wall, due to stentdeployment, was evaluated as a first factor and graded according to themethod of Schwartz et al. (J. Am. Coll. Cardiol 1992; 19:267-274).Likewise, inflammatory reaction at every stent filament site wasexamined (second factor) by searching for inflammatory cells and gradedas well. Appearance of thrombus was evaluated as a third factor andgraded. The mean value of every factor for the 10 samples of each of thetwo stent types was calculated.

Thrombus formation was decreased in the coated stent group. Alsoperi-stent inflammation was decreased in the TiN-coated stent group.

Finally, a morphometric study was carried out on the stented vesselsegments at the time of follow-up after six weeks of implantation. Thestudy was made using a computerized morphometry program (Leitz CBA8000). Measurements of lumen area, lumen inside the internal elasticlamina (=IEL area) and lumen inside the external elastic lamina (=EELarea) were performed on the arterial sites, all in mm2. Neointimalhyperplasia (=IEL area minus lumen area) and area stenosis in % as theratio of hyperplasia to IEL area were derived therefrom. The results areshown in Table 17.

TABLE 17 Morphometry of titanium nitride coated stents Mean CrossSection Non-coated TiN-coated Area (mm²) stent n = 10 stent n = 10 Lumenarea (mm²) 1.71 ± 0.66 2.86 ± 0.74 IEL area (mm²) 3.87 ± 1.39 3.81 ±1.02 EEL area (mm²) 5.74 ± 2.06 5.86 ± 2.12 Hyperplasia (mm²) 2.16 ±1.48 0.95 ± 0.64 Area stenosis (%) 54 ± 15 25 ± 11

TiN-coated stents showed an improved lumen area and a decreasedneointimal hyperplasia and area stenosis at follow-up. Although theinvention has been described for coronary blood vessels, similar resultscan be obtained for stents and intraluminal prosthesis with aTiN-coating in other luminal life stream conducts in animal and humanbodies.

5) A Local Intraluminal Medicine or Gene Releasing System

Several trials with systematically (oral or intravenous) administeredanti restenotic medicines after dilatation of narrowed lumina (forexample of a coronary arterial atherosclerotic narrowing) failed inconsequence of a too limited medicine concentration on the place wherethe medicine has to act and due to the systemic medicine's side effectswhen higher doses are administered. For this reason medicines wereapplied locally, at the place of the organ to be treated. For example inthe treatment of coronary stenoses using special catheters, medicineswere injected into the vessel wall.

Disadvantages of this approach are the limited efficiency of the socalled local treatment (less than 5% of the administered medicinereaches the target organ) and the increased damage to the target organdue to the local drug administration.

Another method is the covering of an endoluminal prosthesis with apolymer coating and the impregnation of the polymer with a medicine. Thedisadvantage of this method is the limited capacity of the coating andthe too fast release of the medicine.

To optimize this system the present inventor have made holes in theendoluminal prosthesis which holes have either a bottom or extendthrough the prosthesis. These holes are filled with the medicineimpregnated polymer before implantation. The holes can vary in size anddensity. The condition is that the radial force of the prosthesis is notaffected. By applying this technique an increment of the local medicinecapacity of the prosthesis with a factor of one hundred and aconsiderable prolongation of the duration of medicine release can beobtained (weeks instead of days). Animal experimental research showed ahundredfold tissue concentration of the medicine in a porcine modelafter implantation of a polymer coated perforated endoluminal prosthesisin coronary arteries.

Furthermore the duration of medicine release was significantly longerand the presence of the medicine in the vascular tissue wassignificantly longer.

Polymeric drug eluting surface coatings have been described to improvestent biocompatibility by locally releasing the drug at the target site(EP-A-0 623 354).

Disadvantages of this system are:

1) the moderate biocompatibility of the polymers used, resulting in anincreased inflammatory reaction,

2) because only very thin polymer layers can be used and the contactarea is large, the drug release using these coated stents is too fastand because only very thin polymer layers can be applied the total doseof drug loaded on the stent to be locally released is limited.

By making holes in the metal structure of the prosthesis, which holesshow either a bottom or extend right through the metal structure of theprosthesis, (FIG. 2) and filling these holes with a drug or a polymercoating containing one or more medicines with anti thrombotic and/oranti restenotic properties, a prosthesis is developed that veryefficiently releases the medicine gradually and puts the medicinedirectly in contact with the damaged tissue. The prosthesis starts tofunction as a reservoir for the medicine, which is gradually releasedafter implantation of the endoluminal prosthesis to carry out itsfunction.

Instead of conventional medicines also genes can be used that code forcertain substances (proteins) having either an anti thrombotic or ananti restenotic action.

Three significant advantages are obtained by using the prosthesisprovided with holes in comparison with the classical polymer coveredprostheses:

1) The total dose of medicine that can be loaded onto the prosthesisincreases with a factor of one hundred to one thousand, depending on thesize and the amount of holes.

2) By making holes showing a bottom, i.e. non-perforating holes, themedicine release can be directed; either towards the tissue surroundingthe lumen or towards the lumen itself.

3) The release time of the medicine becomes much longer (weeks insteadof days).

After having made the prosthesis, the therapeutic agent, i.e. a medicineor genes is to be applied onto the prosthesis, in particular into theholes provided on its surface. This can be done by dipping and/orspraying, after which the therapeutic agent applied next to the holescan optionally be removed. The therapeutic agent can either be appliedas such or as a solution. In a preferred embodiment, it is howevercombined with a polymer which increases the adherence to the prosthesisand which can be used to control the release properties of thetherapeutic agent. In a preferred embodiment there is applied to thebody of a prosthesis and in particular to its tissue-contacting outersurface, a solution which includes a solvent, a polymer dissolved in thesolvent and a therapeutic substance (i.e. a drug) dispersed in thesolvent, and the solvent thereafter is evaporated to leave a drugeluting polymeric substance filling the holes of the prosthesis. Theinclusion of a polymer in intimate contact with a drug filling up theholes of the prosthesis allows the drug to be retained in the prosthesisin a resilient matrix during expansion of the prosthesis and also slowsthe administration of drug following implantation. This method can beused whether the perforated prosthesis has a metallic or polymericsurface. The method is also an extremely simple one since it can beeffected by simply immersing the perforated prosthesis into the solutionor by spraying the solution onto the perforated prosthesis. The amountof drug to be included in the perforated prosthesis can be readilycontrolled by changing the size and the amounts of the holes and/orperforations or by using different drug concentrations and or differentcoating application methods. The rate at which the drug is delivered canbe controlled by the selection of an appropriate bioabsorbable orbiostable polymer and by the ratio of drug to polymer in the solution.By this method, drugs such as glucocorticoids (e.g. methylprednisolone,dexamethasone, betamethasone), heparin, hirudin, tocopherol,angiopeptin, aspirin, ACE inhibitors, Cytochalasin A, B, and D,Trapidil, Paclitaxel, Rapamycin, Actinomycin, growth factors,oligonucleotides, and, more generally, antiplatelet agents,anticoagulant-agents, antimitotic agents, antioxidants, antimetaboliteagents, and anti-inflammatory agents and also genes can be stored in aperforated prosthesis, retained in a perforated prosthesis duringexpansion of the perforated prosthesis and elute the drug at acontrolled rate. The release rate can be further controlled by usingadditional barrier coatings or multiple layers of coating with varyingdrug concentrations. In operation, the perforated prosthesis madeaccording to the present invention can deliver drugs to a body lumen byintroducing the perforated prosthesis transluminally into a selectedportion of the body lumen and radially expanding the perforatedprosthesis into contact with the body lumen. The transluminal deliverycan be accomplished by a catheter designed for the delivery ofperforated prostheses and the radial expansion can be accomplished byballoon expansion of the perforated prosthesis, by self-expansion of theperforated prosthesis or a combination of self-expansion and balloonexpansion.

The underlying structure of the perforated prosthesis used according tothe invention can be virtually any perforated prosthesis design, forexample of the self-expanding type or of the balloon expandable type,and of metal or polymeric material. Thus metal prosthesis designs suchas those disclosed in U.S. Pat. No. 4,733,665 (Palmaz) and U.S. Pat. No.5,603,721 (Lau) could be used in the present invention. The perforatedprosthesis could be made of virtually any biocompatible material havingphysical properties suitable for the design. For example, tantalum,nitinol and stainless steel have been proven suitable for many suchdesigns and could be used in the present invention. Also, prosthesesmade of biostable or bioabsorbable polymers such as poly(ethyleneterephthalate), polyacetal, poly(lactic acid), poly(ethyleneoxide)/poly(butylene terephthalate) copolymer could be used in thepresent invention. Although the perforated prosthesis surface should beclean and free from contaminants that may be introduced duringmanufacturing, the perforated prosthesis surface requires no particularsurface treatment in order to retain the coating applied in the presentinvention.

To coat the perforated prosthesis, in particular to fill the holes madetherein, the following method can be followed. A solution which includesa solvent, a polymer dissolved in the solvent and a therapeuticsubstance dispersed in the solvent is first prepared. The solvent,polymer and therapeutic substance should be mutually compatible. Thesolvent should be capable of placing the polymer into solution at theconcentration desired. Moreover the solvent and polymer should notchemically alter the therapeutic character of the therapeutic substance.However, the therapeutic substance only needs to be dispersed throughoutthe solvent so that it may be either in a true solution with the solventor dispersed in fine particles in the solvent. Examples of some suitablecombinations of polymer, solvent and therapeutic substance are set forthin Table 18.

TABLE 18 Examples of some suitable combinations of polymers, solventsand therapeutic substances Therapeutic Polymer Solvent substancepoly(L-lactic acid) chloroform dexamethasone poly(lactic acid- acetonedexamethasone co-glycolic acid) polyether urethane N-methyl tocopherolpyrrolidone (vitamin E) silicone adhesive xylene dexamethasone phosphatepoly(hydroxybutyrate- Dichloromethane Asprin co-hydroxy-valerate) fibrinWater (buffered heparin saline)

The solution is applied to the perforated prosthesis and the solvent isallowed to evaporate, thereby filling the perforations and leaving onthe perforated prosthesis surface a coating of the polymer and thetherapeutic substance. Typically, the solution can be applied to theperforated prosthesis by either spraying the solution onto theperforated prosthesis or immersing the perforated prosthesis in thesolution. Whether one chooses application by immersion or application byspraying depends principally on the viscosity and surface tension of thesolution. After having coated the prosthesis, the prosthesis canoptionally be cleaned to remove the coating applied next to the holesleaving only the therapeutic agent present in the holes onto theprosthesis.

The polymer chosen should be a polymer that is biocompatible andminimizes irritation to the vessel wall when the perforated prosthesisis implanted. The polymer may be either a biostable or a bioabsorbablepolymer depending on the desired rate of release or the desired degreeof polymer stability, but a bioabsorbable polymer may be more desirablesince, unlike a biostable polymer, it will not be present long afterimplantation to cause any adverse, chronic local response. Bioabsorbablepolymers that could be used include poly(L-lactic acid),polycaprolactone, poly(lactide-co-glycolide), poly(hydroxybutyrate),poly(hydroxyburytate-co-valerate), polydioxanone, polyorthoester,polyanhydride, poly(glycolic acid), poly(D, L-lactic acid),poly(glycolic acid-co-trimethylene carbonate), polyphosphoester,polyphosoester urethane, poly(amino acids), cyanoacrylates,poly(trimethylene carbonate), poly(iminocarbonate), copoly(etheresters)(e.g. PEO/PLA) polyalkylene oxalates, poly(organs)phosphazenes,hydrophylic polymetracrylates and biomolecules such as fibrin,fibrinogen, cellulose, starch, collagen and hyaluronic acid. Also,biostable polymers with a relative low chronic tissue response such aspolyurethanes, silicones, and polyesters could be used and otherpolymers could also be used if they can be dissolved and cured orpolymerized on the perforated prosthesis such as polyolefins,polyisobutylene and ethylene-alphaolefin copolymers; acrylic polymersand copolymers, vinyl halide polymers and copolymers, such as polyvinylchloride; polyvinyl ethers, such as polyvinyl methyl ether;polyvinylidene halides, such as polyvinylidene fluoride andpolyvinylidende chloride; polyacrylonitrile, polyvinyl ketones;polyvinyl aromatics, such as polystyrene, polyvinyl esters, such aspolyvinyl acetate; copolymers of vinyl monomers with each other andolefins, such as ethylene-methyl methacrylate copolymers,actrylonitrile-styrene copolymers, ABS resins, and ethylene-vinylacetate copolymers; polyamides, such as Nylon 66 and polycaprolactam;alkyl resins; polycarbonates: polyoxymthylenes; polyimides; polyethers;epoxy resins, polyurethanes; rayon; rayon-triacetate; cellulose,cellulose acetate, cellulose butyrate; cellulose acetate butyrate;cellophane; cellulose nitrate; cellulose propionate; cellulose ethers,carboxymethyl cellulose and hydrophylic polymetacrylates. The ratio oftherapeutic substance to polymer in the solution will depend on theefficacy of the polymer in securing the therapeutic substance into theperforated prosthesis and the rate at which the coating is to releasethe therapeutic substance to the tissue of the blood vessel or bodyconduit. More polymer may be needed if it has relatively poor efficacyin retaining the therapeutic substance in the perforated prosthesis andmore polymer may be needed in order to provide an elution matrix thatlimits the elution of a very soluble therapeutic substance. A wide ratioof therapeutic substance to polymer could therefore be appropriate andthe weight ratio could range from about 10:1 to 1:100. The therapeuticsubstance could be virtually any therapeutic substance which possessesdesirable therapeutic characteristics for application to a blood vesselor body conduit. This can include both solid substances and liquidsubstances. For example, glucocorticoids (e.g. methyl prednisolone,dexamethasone, betamethasone), heparin, hirudin, tocopherol,angiopeptin, aspirin, Cytochalasin A, B, and D, Trapidil, Paclitaxel,Rapamycin, Actinomycin, ACE inhibitors, A2 blockers, beta blockers,growth factors, oligonucleotides, and, more generally, antiplateletagents, anticoagulant agents, antimitotic agents, antioxidants,antimetabolite agents, and anti-inflammatory agents could be used.Antiplatelet agents can include drugs such as aspirin and dipyridamole.Anticoagulant agents can include drugs such as heparin, coumadin,protamine, hirudin and tick anticoagulant protein. Antimitotic agentsand antimetabolite agents can include drugs such as methotrexate,azathioprine, vincristine, vinblastine, fluorouracil, adriamycin andmitomycin. Furthermore this perforated prosthesis can be used to delivergenes that code for substances that can influence the foreign bodyreaction to the prosthesis or modify the healing response induced bytissue damage.

Description of the Perforated Prosthesis in Order to Obtain an ImprovedLocal Drug Delivery Device

To illustrate the invention a tubular laser cut balloon expandable stentwas used (FIGS. 1 and 2). Perforating holes of 50 μm were made, i.e.holes extending entirely through the stent, using the same water-guidedeximer laser as described in the first part of this description at afrequency of 100 Hz, pulse duration of 0.15 ms and a voltage of 510volts. In this way the holes can easily be made in the same operation asthe cutting of the prosthesis out of the tubular member. Other methodscan however also be used to make the holes, for example mechanical diecutting or conventional laser cutting techniques. FIG. 14 shows amicroscopic picture of a part of the obtained prosthesis. Afterconventional electrochemical polishing the stents were dipped in apolymer solution in which the drug was dissolved. In this example usewas made of a fluorinated polymethacrylate (PFM-P75) in which 10% methylprednisolone was dissolved. Total loading dose of methyl prednisoloneloaded on a PFM-P75-coated non perforated stents was 10 μg. Totalloading dose of a perforated stent was 3500 μg.

In vitro release curves of the methyl prednisolone loaded PFM-P75-coatedstents showed a gradually release of the methyl prednisolone over 3weeks compared to 48 hours for the non perforated stents. Implantationof the methyl prednisolone loaded perforated stents in porcine coronaryarteries using the same study protocol as for the TiN-coated stentsdemonstrated perfect biocompatibility of these stents. No inflammationsurrounding the stent filaments was found at day 3, day 7 and day 14. Atsix weeks only a minimal neointimal hyperplasia was found. Thisinvention can be used with all kinds of drug or gene containing polymersand also for direct coating of drugs or genes onto the prosthesiswithout the use of a polymer.

In general, the holes provided in the prosthesis do not have to extendthrough the prosthesis as in the previous example but show preferably abottom in order to obtain a directional release of the therapeuticagent. Different types of holes are illustrated in the schematicdrawings of FIGS. 9 to 13 and in the microscopic pictures of FIGS. 14 to16. As already explained hereabove, the prosthesis comprises a tubularwall which is usually produced from solid sheet metal but which may alsobe made of a synthetic sheet material. The wall comprises cuts, usuallyaround cut away portions, forming the struts 1 of which the prosthesisis composed. These struts have an outer or convex surface 2, arranged toengage after implantation the inner wall of the lumen, and an oppositeinner or concave surface 3. In order to increase the therapeutic agentloading capacity of the prosthesis, holes 4 are made in the outersurface of the tubular wall. FIGS. 9 to 13 are cross-sectional viewsthrough a strut at the location of such a hole 4. As indicated on thesefigures, the strut has a strut width W and a thickness T, thelongitudinal direction of the strut is indicated by reference A in FIG.2.

At the outer surface of the strut, the holes 4 show an outer opening 5which is situated at a distance from both longitudinal edges 10 of thestrut. This outer opening has a width w measured perpendicular to thelongitudinal direction A of the strut and a length I measured in thislongitudinal direction. Since the edge of the opening 5 is somewhatbeveled, the width and the length of the opening 5 is to be determinedby drawing a line 11 along the inner wall of the hole 4 and determiningthe point of intersection with the outer surface plane so that the bevelof the upper edge of the hole 4 is not taken into account fordetermining the width and length of its outer opening 5. The same goesfor the inner opening which will be described hereinafter in case of aperforating hole.

According to this aspect of the invention, the length I of the outeropening 5 should comprise at the most five times, and preferably at themost three times, the width w thereof whilst the hole 4 itself shouldextend over a depth d in the strut which is larger than 30%, preferablylarger than 50%, and most preferably larger than 60%, of the thickness Tof the strut 1. In this way, the therapeutic agent is distributed over anumber of relatively small holes enabling a homogeneous distributionthereof over the surface of the prosthesis. The total amount oftherapeutic agent applied onto the prosthesis can be controlled not onlyby the number of holes but also by the depth thereof. An advantage ofproviding deeper holes is that the surface of the outer opening 5through which the therapeutic agent can be released is relatively smallcompared to the volume of the hole so that the duration of thetherapeutic agent release can be extended.

The outer openings 5 have advantageously a width w larger than 10 μm, inparticular larger than 20 μm and more particularly larger than 30 μm butsmaller than 100 μm, preferably smaller than 60 μm and most preferablysmaller or equal to 50 μm. The length of the outer openings 5 maycomprise up to five times this width but is preferably substantiallyequal to the width w. The opening 5 is in particular preferablysubstantially circular.

In a preferred embodiment, the width w of the outer opening 5 comprisesat the most 60%, preferably at the most 50%, of the width W of the strut1. Together with the limited length I of the outer openings 5 thisrelatively small width enables to increase the depth d of the holeswhilst maintaining the required minimum radial strength of theprosthesis.

As illustrated in FIG. 2, the openings are divided according to thelongitudinal directions of the struts, i.e. they are not preferably notarranged next to one another in the transverse direction in order tohave a minimal effect on the radial strength of the prosthesis.Preferably, the holes are arranged on substantially constant mutualdistances to achieve a uniform distribution of the therapeutic agent.

FIGS. 9 and 14 illustrate a perforating hole 4 forming at the innersurface 3 of the tubular prosthesis wall an inner opening 6 which issubstantially as large as the opposite outer opening 5. Apart from thebeveled edges, the hole 4 is substantially cylindrical. Such a hole 4can easily be made by means of a laser beam, preferably by means of aliquid-guided laser beam by applying a total amount of cutting energysufficient to produce such a hole without having to move the laser beam.

Referring now to FIGS. 9 and 10, the inner opening 6 has a first sizeand the outer opening 5 has a second size. As seen in both FIGS. 9 and10, the width w of the hole 4 is no greater than at least one of thefirst size or the second size. The first size of the inner opening andthe second size of the outer opening define the maximum width w of thehole 4 over its depth d.

FIGS. 10, 12 and 15 illustrate a perforating hole 4 which also form aninner opening 6 but which has the advantage that there is a moredirectional release of the therapeutic agent, i.e. more therapeuticagent is released towards the wall of the lumen than towards theinterior thereof. This is due to the fact that the inner opening 6 issmaller than the outer opening 5. In FIG. 10, the hole narrows conicallyfrom the outer opening 5 towards the inner opening 6 and is thussubstantially conical. In FIG. 12, the hole shows on the contrary firsta substantially cylindrical portion 7 and subsequently a bottom portion8 narrowing conically towards the inner opening 6. These two types ofholes can also easily be made by means of a laser beam, preferably bymeans of a liquid-guided laser beam, without having to move the laserbeam during the cutting operation. The different types of holes can inparticular be controlled by adjusting the pulse width of the pulse laserbeam, a greater pulse width producing a more elongated cone shape as inFIG. 10 whilst a smaller pulse width produces a shorter or steeper coneshape as in FIG. 12.

In the most advantageous embodiment, the total amount of energy of thelaser beam is reduced so that the hole does not extend entirely throughthe strut but forms a bottom 9. Such a hole is illustrated in FIGS. 11,13 and 16. In the illustrated embodiments, the hole is entirely conicalor frusto-conical. It will however be clear that the hole could alsoshow a conical bottom part and on top of that a more cylindrical part,depending on the thickness of the strut and the depth of the hole. Fromthe shapes of the conical portions in the figures it will be clear thatthe hole illustrated in FIG. 11 can be achieved by means of a same typeof laser beam as the hole illustrated in FIG. 10 whilst the holeillustrated in FIG. 13 can be achieved by means of the same type oflaser beam as the hole illustrated in FIG. 12, the total amount ofcutting energy, i.e. the duration of the cutting process being of coursereduced to achieve a shallower hole.

The holes illustrated in the figures have all a substantially circularcross-section seen parallel to the outer or inner surface of theprosthesis. For making elongated holes, i.e. holes having a length Ilarger than their width w, two or more of the above described holes canbe made next to one another to achieve one elongated hole.

The above disclosure is intended to be illustrative and not exhaustive.This description will suggest many variations and alternatives to one ofordinary skill in this art. The various elements shown in the individualfigures and described above may be combined or modified for combinationas desired. All these alternatives and variations are intended to beincluded within the scope of the claims where the term “comprising”means “including, but not limited to”.

Further, the particular features presented in the dependent claims canbe combined with each other in other manners within the scope of theinvention such that the invention should be recognized as alsospecifically directed to other embodiments having any other possiblecombination of the features of the dependent claims. For instance, forpurposes of claim publication, any dependent claim which follows shouldbe taken as alternatively written in a multiple dependent form from allprior claims which possess all antecedents referenced in such dependentclaim if such multiple dependent format is an accepted format within thejurisdiction (e.g. each claim depending directly from claim 1 should bealternatively taken as depending from all previous claims). Injurisdictions where multiple dependent claim formats are restricted, thefollowing dependent claims should each be also taken as alternativelywritten in each singly dependent claim format which creates a dependencyfrom a prior antecedent-possessing claim other than the specific claimlisted in such dependent claim below.

This completes the description of the preferred and alternateembodiments of the invention. Those skilled in the art may recognizeother equivalents to the specific embodiment described herein whichequivalents are intended to be encompassed by the claims attachedhereto.

1. A radially expandable prosthesis comprising: a tubular wallcomprising a plurality of adjacent filaments, each filament having aplurality of adjacent segments, two adjacent segments defining a segmentpair, segments of segment pairs having connected ends and unconnectedends, the connected ends connected to one another via a curved section,the curved section comprising a proximal turn or a distal turn, eachsegment having an inflection point along its length, wherein thesegments of the segment pairs are separated from one another, in acircumferential direction, by a greater distance at their connected endsthan at their unconnected ends; a plurality of axial elements joiningthe adjacent filaments, each axial element consisting of a singlecontinuous curve; a plurality of reservoirs disposed within the adjacentfilaments; and a titanium nitride coating disposed on the tubular wall,wherein the titanium nitride coating has a thickness between 1 to 10 μm.2. The prosthesis of claim 1, wherein the titanium nitride coating isporous.
 3. The prosthesis of claim 1, wherein the reservoirs are holesextending through the entire thickness of the adjacent filaments.
 4. Theprosthesis of claim 1, wherein a plurality of reservoirs are disposedonly within the adjacent filaments and there are no reservoirs in theconnectors.
 5. The prosthesis of claim 1, wherein each filament has anouter surface, an inner surface, and a thickness therebetween, and eachreservoir comprises an opening in an outer surface of the strut thatextends through only a portion of the thickness of the filament.
 6. Aradially expandable prosthesis comprising: a tubular wall having anouter surface, an inner surface, and a thickness therebetween, thetubular wall comprising a plurality of interconnected serpentine bandsincluding first bands and second bands, each band comprising adjacentsegments, two adjacent segments defining a segment pair, segments ofsegment pairs having connected ends and unconnected ends, the connectedends connected to one another via a curved section, the curved sectioncomprising a proximal turn or a distal turn, each segment having aninflection point along its length, wherein the segments of the segmentpairs are separated from one another, in a circumferential direction, bya greater distance at their connected ends than at their unconnectedends; the bands connected one to another via a plurality of connectors,each connector consisting of a single continuous curve having a firstend and a second end, wherein the first end is circumferentially alignedwith the second end along a common longitude; the first bands beingimmediately adjacent to two second bands, the second bands beingimmediately adjacent to two first bands, each first band being out ofphase with each second band; the connectors extending between distalvalleys and proximal peaks of immediately adjacent bands; every thirddistal valley of each of the first and second serpentine bands havingone of the connectors attached thereto; the serpentine bands comprisinga plurality of reservoirs; wherein the prosthesis has anelectrochemically polished surface.
 7. The prosthesis of claim 6,wherein the reservoirs are holes extending from the outer surface of thetubular wall, the reservoirs having a depth of at least 30% of thethickness of the tubular wall.
 8. The prosthesis of claim 7, wherein thereservoirs have a depth of 50% of the thickness of the tubular wall. 9.The prosthesis of claim 7, wherein the reservoirs have a depth of 60% ofthe thickness of the tubular wall.
 10. The prosthesis of claim 7,wherein the reservoirs extend through the entire thickness of thetubular wall.
 11. The prosthesis of claim 6, wherein the reservoirs arefilled with a therapeutic agent.
 12. The prosthesis of claim 11, whereinat least one of the outer surface and the inner surface of the tubularwall is coated with a therapeutic agent.
 13. The prosthesis of claim 6further comprising a titanium nitride coating, wherein the titaniumnitride coating has a thickness between 0.1 to 500 μm.
 14. Theprosthesis of claim 6, wherein the single continuous curve of each ofthe connectors has a connector peak.
 15. The prosthesis of claim 14,wherein each of the connector peaks has the same orientation.
 16. Theprosthesis of claim 6, wherein a plurality of reservoirs are disposedonly within the serpentine bands and there are no reservoirs in theconnectors.
 17. The prosthesis of claim 6, wherein the connectorsconnect one distal valley of the first band to a confronting proximalvalley of the second band.